Computed tomography (CT) is a well-known method for obtaining images of the internal structure of a subject from its projections. In X-ray CT, these projections are obtained by rotating an external radiation source, usually an X-ray tube, around the subject and measuring the X-ray transmission through the field of view with an opposite X-ray detector array. The measured intensity of the transmitted X-ray flux through the subject, in reference to the measured intensity without the subject, provides a measure of the mean attenuation through the body tissues, which in turn provides information on the tissue density and composition. A set of projections, obtained over 180° or 360°, is then processed by a tomographic reconstruction algorithm which creates the cross sectional image.
CT has been shown to be superior to conventional radiography in the detection of a wide variety of diseases because of the greater contrast it allows to achieve. However, CT involves a considerably greater amount of radiation than conventional radiography or other tomographic imaging modalities such as Single Photon Emission Computed Tomography (SPECT) and Positron Emission Tomography (PET).
Modern CT scanners are used more and more for routine clinical check-up and they currently account for 30% to 50% of the radiation dose to the general population. In typical clinical CT diagnostic investigations, the radiation dose received by the patient typically exceeds the annual dose limit allowed to the normal population. This issue is particularly preoccupying in pediatric scanning, as the human body is still growing and more prone to developing cancer 10 or 20 years later as a result of dose exposure.
For repeated studies in the same subject, as it is required for example for preventive screening or patient follow-up, it is of utmost importance to reduce the radiation exposure to the minimum level compatible with the diagnostic application. However, reducing the radiation dose on CT procedures has the effect to increase image noise, thus reducing contrast, making lesions more difficult to detect and ultimately affecting diagnostic accuracy.
In micro-CT imaging of animal subjects, the typical radiation dose received in one single scan can represent as much as 10% of the LD50/30 for mice (the dose required to kill 50% of mice by 30 days after radiation exposure without other intervention). Significant short-term stimulation effects (DNA repair mechanisms, immune response, free-radical detoxification and apoptosis) and long-term effects of radiation-induced damage have been reported in this dose range, which may potentially have confounding biological effects biasing research results.
Current CT scanners employ ionization gas detectors (e.g., xenon), semiconductor diodes (Si, CdTe, CZT . . . ), phosphors coupled to charge-coupled devices (CCDs), or scintillators coupled to silicon diodes or photomultiplier tubes. Due to limitations in signal-to-noise ratio and/or count rate, these detectors must be operated in current mode, whereby the product of the mean X-ray event rate and the average X-ray energy is the measured parameter. Also depending on the material used for X-ray detection, the detector's quantum efficiency is sometime well below the ideal 100% value. As a consequence, no energy dependent processing (such as multi-spectral image analysis or scatter correction) may be performed. In spite of the fact that current mode CT involves a poor utilization of the information conveyed by the number and energy of the individual transmitted X-rays, it is well suited for high rate studies where high X-ray fluxes and fast scanning times are employed.
In another respect, the inherent integration of the X-ray beam energy has the detrimental consequence of exacerbating the so-called beam hardening effect, by increasing the weight of high energy X-rays relative to low energy X-rays proportionally to the X-ray photon energy. However, transmitted low energy photons convey more contrast information about soft tissue than transmitted photons at high energy. The ideal weight factor to achieve maximum contrast using spectral X-ray sources is proportional to E−3 (where E is the incident photon energy) to reflect the attenuation properties of materials which follows the following equation in the diagnostic energy range [1]:
      μ    ⁡          (      E      )        ≈                              N          0                ⁢        ρ            A        ⁢          (                                    aZ            4.2                    ⁢                      E                          -              3                                      +        bZ            )      where N0 is Avogadro's number, ρ is the density, A is the atomic mass and Z is the atomic number. For integrating systems, which inherently take weight factor to be proportional to the photon energy, there is a difference of the order of E4 relative to optimum weighting.
Another adverse consequence of X-ray integrating systems, which imposes strict stability requirements on the entire systems, is that the noise from all sources (electronic, variance due to scintillation photon or charge carrier statistics, afterglow in phosphors, systematic signal bias) is integrated and measured together with the signal in acquiring CT data, leading to noisier projection data and degraded image contrast. As a consequence, higher doses are required to overcome the intrinsic noise in the signal and to achieve the required contrast in the CT images. Another related consequence is that more powerful, cumbersome and expensive X-ray tubes with complex cooling systems must be used to reduce imaging times.
Single photon counting systems have been developed in some other imaging applications than CT, such as conventional scintigraphy, SPECT and PET. However, the signals from the detectors in most of these imaging systems are multiplexed or combined together in order to process the signals from a large number of detector elements (or pixels) using a smaller number of electronic channels. On one hand, the signal amplitude generated by the detectors in these applications (incident radiation>100 keV) is generally sufficient to allow sufficiently accurate computation of the position of interaction. On the other hand, this approach is advantageous to reduce the cost and complexity of the systems, but it greatly limits the maximum count rate per detector element that can be achieved, which is well below the mean count rate per unit area required in CT imaging.
Strip detector configuration made of semiconductor materials has been proposed to measure low energy radiation in the diagnostic X-ray range [2]. Such detectors, made of CdTe, CZT and Si, can operate at room temperature and provide adequate signal to noise ratio to measure the energy of individual X-ray photons with high accuracy. However, the multiplexing of N2 detector pixels into 2N electronic channels reduces the maximum count rate per pixel by a factor of at least N/2 (neglecting the time required for decoding). It has been found that such a system is severely statistics limited for high-rate photon counting CT. The use of an individual readout channel per detector pixel was also found to be count rate limited due to the long charge collection time which increases dead time and severely limits the maximum event rate that can be processed [3]. Semiconductor detectors also suffer from low detection efficiency in the higher diagnostic energy range. Even though thicker semiconductor diodes has been proposed to overcome this problem, their use increases the cost of the detector and the charge migration time adds extra dead time that further limits the detector count rate.
Pixelated detectors made of high purity germanium (HPGe) have the advantage with respect to energy resolution of being capable of resolving the fine structure of X-ray spectra. However, such systems require detector cooling, typically to 77° K, and they must be used in conjunction with low-noise charge preamplifiers having a long integration time to collect the slow drifting charge carriers from the bulk of the detector material. Either in strip detector configuration or with an individual readout channel per detector pixel, pulse pile-up of the slow decaying signals from the charge preamplifiers limits the maximum count rate that can be reached with such detectors. Moreover, such systems are generally much too expensive to be considered for a large-scale application, such as medical imaging scanners.
Yet another method from the prior art for counting and measuring energy of individual X-ray photons is to read out detector pixels at a rate such that the likelihood of registering more than one X-ray photon per detector pixel during a readout period is negligible. The readout circuit is made of register cells and a controller to transfer the response in pixels to register readout cells. If the response of the detector pixels may be weighted according to the energy of the detected X-ray photon, a detection mechanism can be implemented for converting the response of the detector pixel into an electric signal (charge or current) that is proportional to the energy of the detected X-ray photon, assuming that the likelihood of arrival of more than one photon in the detector pixel during one readout period is negligible. A drawback of this method is that it requires ultra-low noise detectors and extremely fast readout rates to achieve the detector count rate required for use in CT imaging.